Membrane and electrode structure for implantable sensor

ABSTRACT

The invention is the design of a biological measuring device for the determination of the concentration of biomolecules (e.g. glucose) in an environment which is designed for implantation into an individual or for use in the context of an external apparatus. The device contains a composite membrane that is essentially entirely permeable to oxygen and permeable to larger biomolecules only in discrete hydrophilic regions. The membrane diffusionally limits the access of biomolecules to an enzyme, present in the hydrophilic region that catalyzes the oxidation of the biomolecule to produce hydrogen peroxide. A sensor in communication with the hydrophilic region is used to determine the amount of product produced or the amount of excess oxygen present allowing for the concentration of the biomolecule to be determined.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application claims the benefit of priority of U.S. provisionalapplication Serial No. 60/269,169 filed Feb. 15, 2001 which isincorporated herein by reference in its entirety.

STATEMENT OF GOVERNMENT INTEREST

The invention was made with Government support under Grant No. DK55064awarded by the National Institutes of Health.

FIELD OF THE INVENTION

The invention relates to the design and use of a biological measuringdevice containing a novel membrane structure.

BACKGROUND OF THE INVENTION

It is standard practice to treat diabetes mellitus predominantly withinsulin injections to compensate for the inability of the pancreas tomake insulin to regulate blood glucose levels. The more tightly a personwith diabetes is able to regulate his or her blood sugar, the lessdetrimental the disease is to overall health. The regulation of bloodglucose would benefit from a glucose sensing device implanted in thebody to monitor blood glucose levels at more frequent intervals than canbe done with presently available repeated blood sampling.

A variety of biomedical measuring devices are routinely used byphysicians and clinicians to monitor physiological variables such asrespiratory rate, blood pressure and temperature. In addition to therepertoire of devices listed above is the enzyme electrode. Enzymeelectrodes enable the user to determine the concentration of certainbiochemicals rapidly and with considerable accuracy by catalyzing thereaction of a biochemical and a detectable coreactant or producing aproduct that may be readily sensed by well-known electrodes (e.g.oxygen, H₂O₂). Currently there are enzyme electrodes that can detecturea, uric acid, glucose, various alcohols, and a number of amino acidswhen used in certain well-defined situations.

A number of variations of the glucose enzyme electrode have beendeveloped, all based on the same reaction catalyzed by glucose oxidase.

To accurately measure the amount of glucose present, both oxygen andwater must be present in excess. As glucose and oxygen diffuse into animmobilized membrane phase, the glucose reacts with oxygen and water toproduce H₂O₂ (hydrogen peroxide). Glucose is detected electrochemicallyusing the immobilized enzyme glucose oxidase coupled to an oxygen- orhydrogen peroxide-sensitive electrode. The reaction results in areduction in oxygen and the production of hydrogen peroxide proportionalto the concentration of glucose in the sample medium.

The electrode can be polarized cathodically to detect residual oxygennot consumed by the enzymatic process, or polarized anodically to detectthe product of the enzyme reaction, hydrogen peroxide. A functionaldevice is composed of at least two detecting electrodes, or at least onedetecting electrode and a reference signal source, to sense theconcentration of oxygen or hydrogen peroxide in the presence and absenceof enzyme reaction. Additionally, the complete device contains anelectronic control means for determining the difference in theconcentration of the substances of interest. From this difference, theconcentration of glucose can be determined.

The enzyme catalase may be included in the oxygen-based system in excessin the immobilized-enzyme phase containing the glucose oxidase tocatalyze the following reaction:

Hence, the overall reaction becomes:

glucose+½O₂→gluconic acid.

This mixture of immobilized enzymes can be used in the oxygen-baseddevice, but not the peroxide-based device. Catalase prevents theaccumulation of hydrogen peroxide which can promote the generation ofoxygen free radicals that are detrimental to health.

Glucose measuring devices for testing of glucose levels in vitro basedon this reaction have been described previously (e.g. Hicks et al., U.S.Pat. No. 3,542,662) and work satisfactorily as neither oxygen nor waterare severely limiting to the reaction when employed in vitro.Additionally, a number of patents have described implantable glucosemeasuring devices. However, certain such devices for implantation havebeen limited in their effectiveness due to the relative deficit ofoxygen compared to glucose in tissues or the blood stream (1: 50-1000).

Previous devices (e.g. Fisher and Abel) have been designed such that thesurface of the device is predominantly permeable to oxygen, but notglucose, and is in contact with the enzyme layer. Glucose reaches theenzyme layer through a minute hole in the oxygen-permeable outer layerthat is in alignment with an electrode sensor beneath it. Hydrogenperoxide produced by the enzyme reaction must diffuse directly to thesensing anode or through a porous membrane adjacent to the electrode,but is otherwise substantially confined within the enzyme layer by theoxygen-permeable layer resulting in unavoidable peroxide-mediated enzymeinactivation and reduced sensor lifetime.

The strategy of designing devices with differentially permeable surfaceareas to limit the amount of glucose entering the device, whilemaximizing the availability of oxygen to the reaction site, is nowcommon (Gough, U.S. Pat, No. 4,484,987). An example based on devicegeometry is seen in Gough, U.S. Pat, No. 4,671,288, which describes acylindrical device permeable to glucose only at the end, and with boththe curved surface and end permeable to oxygen. Such a device is placedin an artery or vein to measure blood glucose. In vascular applications,the advantage is direct access to blood glucose, leading to a relativelyrapid response. The major disadvantage of vascular implantation is thepossibility of eliciting blood clots or vascular wall damage. Thisdevice is not ideal for implantation in tissues.

An alternative geometrically restricted device assembly was described inGough, U.S. Pat, No. 4,650,547. The patent teaches a “stratified”structure in which the electrode was first overlaid with anenzyme-containing layer, and second with a non-glucose-permeablemembrane. The resulting device is permeable to oxygen over the entiresurface of the membrane. However, glucose may only reach the enzymethrough the “edge” of the device in a direction perpendicular to theelectrode, thus regulating the ratio of the access of the two reactantsto the enzyme.

Devices have been developed for implantation in tissue to overcomepotential problems of safely inserting into, and operating sensorswithin, the circulatory system (e.g. Gough, U.S. Pat. No. 4,671,288);however, their accuracy may be limited by the lower availability ofoxygen in tissues. The device membrane is a combination ofglucose-permeable area and oxygen-permeable domains. The ratio of theoxygen-permeable areas to the glucose-permeable areas is somewhatlimited due to the design.

To avoid geometric restrictions on devices, membranes that are variablypermeable to oxygen and glucose have been developed (Allen, U.S. Pat.No. 5,322,063). Membrane compositions are taught in which the relativepermeability of oxygen and glucose are manipulated by altering the watercontent of a polymeric formulation. The disadvantages of such a membranemay include sensitivity of the membrane performance to variables duringmanufacture and that regions of oxygen permeability may not be focusedover electrodes within the device.

An alternative strategy to device construction is to incorporate anenzyme-containing membrane that is hydrophilic and also contains smallhydrophobic domains to increase gas solubility, giving rise todifferential permeability of the polar and gaseous reactants (e.g.Gough, U.S. Pat. Nos. 4,484,987 and 4,890,620). Such membranes readilyallow for the diffusion of small apolar molecules, such as oxygen, whilelimiting the diffusion of larger polar molecules, such as glucose. Thedisadvantage is that the amount of hydrophobic polymer phase must berelatively large to allow for adequate oxygen permeability, therebyreducing the hydrophilic volume available for enzyme inclusionsufficient to counter inactivation during long-term operation.

Schulman et al. (U.S. Pat. No. 5,660,163) teach a device with a siliconerubber membrane containing at least one “pocket” filled with glucoseoxidase in a gelatinous conductive solution located over a first workingelectrode. In a preferred embodiment, the length of the “pocket” isapproximately 3 times its thickness to optimize the linearity betweencurrent and the glucose concentration measurement. Because the long axisof the “pocket” is oriented parallel to the electrode surface, thisdesign may be less amenable to miniaturization for tissue implantation.

SUMMARY OF THE INVENTION

The invention is the design and use of a biological measuring device forimplantation into an individual or for use in an external environment.The device contains an enzyme electrode to detect the coreactant orproduct (e.g. oxygen, H₂O₂, respectively) of an enzymatic reactioncatalyzed by an oxidase (e.g. glucose oxidase, lactate oxidase,cholesterol oxidase) of the biological molecule of interest (e.g.glucose, lactate, cholesterol) with a limiting reagent or coreactant(e.g. oxygen). The device contains a differentially permeable membranethat limits the access of the biological molecule of interest, which ispresent in the device's environment at a relatively high concentrationas compared to the coreactant, to the enzyme. (Expected ratios ofbiological molecule to coreactant concentrations (e.g. glucoseconcentration to oxygen concentration) in biological samples orenvironments may be expected to range up to 10:1 and beyond, expressedin units of mg/dl/mmHg.) Thus, the biological molecule becomes thelimiting reagent in a critical zone within the enzyme-containing regionof the membrane, allowing for its quantification by assaying the amountof product produced or the amount of unconsumed coreactant by means ofan associated sensor or electrode, responsive to the coreactant orproduct.

The membrane is composed of a continuous or nearly continuousrestricted-permeability membrane body, permeable to oxygen andessentially impermeable to larger biological molecules (e.g. glucose,lactate, cholesterol), and discrete hydrophilic regions, permeable toboth biological molecules and oxygen (FIG. 1). The reactants diffusefrom the environment into the device through a single surface of thedevice. The size, density, shape, and number of hydrophilic regions maybe varied depending upon the bodily fluid, tissue, or environment intowhich the device is implanted or depending upon the choice of theassociated sensor. As opposed to prior membranes which haverestricted-permeability and hydrophilic surfaces at restricted locationson the device defined by device shape, or other sensors covered inmembranes whose differential oxygen- and biologicalmolecule-permeability is continuous, the location, number, shape, andsize of the oxygen- and biological molecule-permeable regions may bemodified to optimize the performance of the sensor.

The invention is a biological measuring device containing the compositemembrane of the invention. The membrane of the invention can beoptimized for detection of a number of biochemicals with a single or aplurality of detecting electrodes. Electrodes may be linked in any of anumber of ways well known to those skilled in the art (e.g. Sargent andGough, 1991, herein incorporated by reference). The size, shape, number,and location of the hydrophilic regions can be varied to deliver theappropriate ranges of the biological molecule and oxygen to the enzymesuch that a detectable amount of product or consumed coreactant reachesthe associated sensor.

The invention is a method to specify the optimal ratio ofrestricted-permeability membrane body to hydrophilic regions in themembrane, and to determine the optimal shape and arrangement of thehydrophilic regions in the membrane such that the concentrations of thereactants in the critical zone are limited by diffusion. Using themethod of the invention, the sensor can be optimized for differentreactions and enzymes for use in different tissues, bodily fluids or inan external sensor.

The invention is the use of the biological measuring device to monitorthe level of a biological molecule, either by implantation in anindividual or by use of the device in an external environment. In apreferred embodiment, the device is used to monitor glucose levels in anindividual with diabetes.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention will be better understood from the followingdetailed description of an exemplary embodiment of the invention, takenin conjunction with the accompanying drawings in which like referencenumerals refer to like parts.

FIG. 1. Schematic of a biological sensor device membrane with a singlehydrophilic region and sensor. The device comprises a membrane body (2)that communicates with an environment (4) having a first material suchas glucose and a second material such as oxygen. The surface (8) of themembrane body communicates with the environment (4). The membrane bodycommunicates with a hydrophilic region (6) with catalyst. Thehydrophilic region (6) contains a critical zone (12) of averageequivalent radius a (14) and length/(16) such that a</. A sensor (18),with surface (20), is sensitive to the reaction product or residualco-reactant and produces a signal in proportion to the concentration. Acontrol (26) responds to the signal, for comparison with a reference(28). The diffusion paths of the first material (30) and (34) and of thesecond material (32) enter the device through the same surface (8). R₁is the radius of the hydrophilic region on the face of the membrane incommunication with the environment. R₂ is the radius of the hydrophilicregion on the face of the membrane in communication with the sensor. R₃is the radius of the sensor.

FIG. 2. Schematic of a biological sensor device for implantation. Thesensor device for implantation comprises a membrane body (2) thatcommunicates with an environment (4) having a first material such asglucose and a second material such as oxygen. The surface (8) of themembrane body communicates with the outer layer (10) to the environment(4). The membrane body communicates with a hydrophilic region (6) withcatalyst. The hydrophilic region (6) contains a critical zone (12) ofaverage equivalent radius a (14) and length/(16) such that a</. A sensor(18), with surface (20), is sensitive to the reaction product orresidual co-reactant and produces a signal in proportion to theconcentration. The sensor's surface (20) communicates with theelectrolyte layer (24) adjacent to the sensor protective layer (22). Acontrol (26) responds to the signal, for comparison with a reference(28). The diffusion paths of the first material (30) and (34) and of thesecond material (32) enter the device through the same surface (8).

FIG. 3. Schematic of a biological sensor device with an alternativehydrophilic region. A biological sensor device having a hydrophilicregion with catalyst with a cross-section in the form of an inverted“T”. The device comprises a membrane body (2) that communicates with anenvironment having a first material such as glucose and a secondmaterial such as oxygen. The surface (8) of the membrane bodycommunicates with the environment. The membrane body communicates with ahydrophilic region (6) with catalyst. The hydrophilic region (6)contains a critical zone (12) of average equivalent radius a andlength/such that a</. A sensor (18), with surface (20), is sensitive tothe reaction product or residual co-reactant and produces a signal inproportion to the concentration. A control (26) responds to the signal,for comparison with a reference (28). The diffusion paths of the firstmaterial (30) and (34) and of the second material (32) enter the devicethrough the same surface (8). L1 is the length of the narrowercylindrical portion of the hydrophilic region and L2 is the full lengthof the hydrophilic region. In this representation, L2 is equivalent toT1 as shown in FIGS. 1 and 4.

FIG. 4. Schematic of membrane illustrating a plurality of hydrophilicregions. This figure describes a sensor device with a membrane body (2)with a plurality of hydrophilic regions (6) with catalyst variouslyjuxtaposed across the sensor surface (20) in communication with thesensor (18), the hydrophilic regions having respective critical zones(12). The diffusion paths for the first material (30) and (34), andsecond material (32) enter the device through the same surface (8). Thesensor (18) with the surface (20) is sensitive to the reaction productor residual co-reactant and produces a signal in proportion to theconcentration. The center-to-center spacing (S) and the radius (R) ofthe hydrophilic regions is shown.

FIG. 5. Schematic of a membrane illustrating a funnel shaped hydrophilicregion as discussed in Example 1 with various specification measurementsindicated, including membrane thickness, enzyme region diameter atenvironment, enzyme region diameter at sensor and height of cylindricalportion of funnel.

FIG. 6. Schematic of a membrane illustrating multiple cylindricalhydrophilic regions as discussed in Example 2 with various specificationmeasurements indicated, including membrane thickness, enzyme regiondiameter and enzyme region spacing.

FIG. 7. The calculated response of an oxygen sensor, in communicationwith hydrophilic regions, to environmental concentrations of glucose andoxygen for various membrane constructions. The electrode current iscalculated and shown as i_(g)/I_(o) which is the ratio of the glucosemodulated oxygen current to the current in the absence of glucose.

DETAILED DESCRIPTION AND PREFERRED EMBODIMENTS

Definitions

Enzymatic sensor assembly—An electrochemical detector component,comprising a noble metal working electrode polarizable as an anode or acathode, potential reference electrode, a counter electrode and layer ofconductive electrolyte forming a thin conductive layer among theelectrode sensor structures;

an electronic polarization and amplification component consisting of apotentiostat or polarizing amplifier, current recording amplifier and asignal conveyor (e.g. a wire); and

a layered or stratified membrane structure composed (1) in the case ofthe oxygen based sensor of an inner, electrode protective layer of apore-free, oxygen-permeable material such as polydimethylsiloxane thatis impermeable to polar compounds, or in the case of a peroxide-basedsensor, a porous membrane that is permeable to hydrogen peroxide andless permeable to larger polar molecules; (2) an enzyme region or domainof specified shape and volume containing immobilized enzymes; (3) amembrane structure for differential control of reactant access to theenzyme region by means of a specified pore size, differentialpermeability reactant solubility or geometric configuration; and (4) anoptional biocompatibility membrane or layer to promote development of abiocompatible interface between tissue or blood and the implanted sensor(FIG. 2). A number of such assemblies are well known such as thosetaught in Schulman, U.S. Pat. No. 5,660,163.

Membrane body—A nearly continuous membrane that is permeable to oxygenand essentially non-permeable to larger biological molecules (e.g.glucose). It may or may not be water-containing and can be made of anyof a number of oxygen-permeable polymeric materials including, but notlimited to, any of the family of silicone-containing,ethylene-containing and propylene-containing polymers, with and withoutfluorine, such as silicone rubbers, polyethylene, polypropylene,Teflons, polyfluorinated hydrocarbons or similar polymers, as well ascertain hydrophilic polymers, such as polyhydroxyethlymethacrylate oflimited molecular porosity, that are permeable to oxygen by virtue ofhaving significant oxygen solubility or diffusivity. Co-polymers,blends, or composites that incorporate these types of materials are alsosuitable.

Hydrophilic region—An intermittent volume in communication with themembrane body that is permeable to both larger biological molecules(e.g. glucose) and oxygen. It can be made of any of a number of glucose-and oxygen-permeable materials including, but not limited to,polyacrylamide gels, glutaraldehyde cross-linked proteins, particularlycollagen or albumin, vinyl pyrollidone, alginates, ethylene oxide,polyhydroxyethylmethacrylate and its derivatives, and other hydrophilicpolymers and co-polymers. Co-polymers, blends, or composites thatincorporate these types of materials are also suitable. An enzyme orcatalyst is typically incorporated into this region.

Critical zone—A volume of the membrane that is coincident with thehydrophilic region, or a portion thereof, through which the reaction ofthe biological molecule with the oxygen is modulated by limiting thediffusion of the biological molecule from the environment. Preferably,it is a volume that is coincident with a given hydrophilic region, or aportion thereof that is bound between two end planes that are orientedperpendicular to the average vector direction of diffusion of thebiological molecule (e.g. glucose) throughout the whole givenhydrophilic region, wherein the average vector direction of diffusion ofthe biological molecule in the critical zone is essentially parallel tothe average vector direction of diffusion of the biological molecule inthe whole given hydrophilic region. Additionally, a critical zone musthave an average equivalent radius that is less than the length of thecritical zone. An equivalent radius is obtained by first dividing by pithe area of a given cross section of a given hydrophilic region, thearea being oriented perpendicular to the average vector direction ofdiffusion of the biological molecule throughout the whole givenhydrophilic region, then taking the square root of the resultingquantity.

Detailed Description

The invention is a novel membrane structure based on a nearly continuousoxygen-permeable, glucose-impermeable membrane body having discreteregions of hydrophilic, glucose-permeable gel in which the enzyme isimmobilized. Additionally, the hydrophilic regions communicate throughthe membrane to one or more underlying electrode sensor structures. Thematerials and methods used for preparing the hydrophilic regions aredescribed in Gough, U.S. Pat. No. 4,484,987 which is incorporated hereinby reference.

The desired geometric relationships between the membrane body and thehydrophilic regions and the shape of the hydrophilic regions mustfunction to supply coreactant to the enzyme gel such that the reactionwithin the gel is limited by the availability of biological moleculerather than coreactant. Any portion of the hydrophilic region that meetsthe definition of critical zone may provide this function. Thehydrophilic regions may or may not penetrate the entire thickness of themembrane, but must communicate, either directly or by means of anexternal membrane having permeability to glucose, with the environmentin which the device is operated. In a preferred embodiment, the deviceis a flat, disc shape. The glucose and oxygen diffuse into the devicethrough a single face at the device-environment interface.

The hydrophilic regions may be varied in size, shape, number and spatialdistribution to advantage in a given device design. Shapes mayinclude: 1) a cylinder orthogonal to the plane of the membrane toprovide radially uniform oxygen access within the enzyme region, 2) asquare or parallelogram, as seen from the face of the membrane, for easeof fabrication by a method of laying one sheet of hydrophobic stripsover another, 3) a cone or other shape of tapering radius, as seen fromthe edge of the membrane with the base at the sensor electrode side toprovide a mechanical confinement of the gel and prevent gel extrusion orseparation from the membrane body during fabrication or useconformations formed from a combination of such shapes, such as a“funnel,” formed by the combination of conically- andcylindrically-shaped regions (e.g. FIGS. 2-3). The exact conformation ofthe shapes listed above is not required.

The size, shape, number, and spatial distribution of the hydrophilic gelregions can be varied (e.g. FIG. 4). The exact patterning of thehydrophilic gel regions is designed to optimize sensor response,sensitivity to biologic molecule, coreactant independence andinsensitivity to environmental heterogeneity. The size of thehydrophilic regions can be varied over different electrodes to providethe sensor with a broader range of sensitivity. It is not necessary forthe sensor to be of the same radius as the hydrophilic region. Moreover,it is possible to design a device with multiple sensors associated witha single hydrophilic region, or multiple hydrophilic regions associatedwith a single sensor. Design choices are based on a variety of factors,such as preference for a particular manufacturing technique,requirements for signal magnitudes based on choice of electroniccircuitry, and the vascular density in the tissue of implantation.

The thickness of the membrane can be controlled to optimize the oxygenindependence, diffusional length for glucose within the hydrophilic gelto provide reserve enzyme, and to optimize respective response times toglucose and oxygen changes.

Regions of the membrane body that can be used to house hydrophilicregions may be fabricated by any of a number of methods well known tothose skilled in the art including programmed laser ablation, molding,cutting, punching, etc. Holes can then be filled with uncrosslinkedenzyme-containing precursor solutions and then crosslinker is added oractivated, to solidify the solution.

A hydrophobic membrane, shown in FIG. 2, may be inserted between theabove-described membrane structure and the oxygen sensing electrode, ordirectly overlying the oxygen electrode and electrolyte solution. Suchan intervening membrane protects the oxygen electrode fromelectrochemical poisoning from polar and diffusable compounds. Itsdimensions and material properties can also be varied to advantagedepending on the exact sensor design. Preferably such a membrane wouldreadily allow the diffusion of oxygen while preventing the diffusion oflarger molecules through the membrane. Additionally, the membrane isthin to maximize the sensitivity of the system to glucose.

The positioning and arrangement of the hydrophilic gel regions can bevaried with regard to the underlying oxygen sensor electrode orelectrodes to optimize the sensitivity and range of the device. It isimportant to note that the sensitivity and response time of the devicecan be altered simply by varying the amount of electrode surface area ofthe oxygen sensor, along with the thickness of the membrane over thesensor. The methods for making these adjustments are well known to thoseskilled in the art.

A number of electrodes and electrode combinations are well known tothose skilled in the art and could be used in this invention. Forexample, the electrodes may be either oxygen or hydrogen peroxidesensing. The sensor may be an electrically conductive layer or anelectrode connected by a wire to single or multichannel electronics.Alternatively, the membrane may be connected directly to theelectronics.

In embodiments of the invention for implantation into the body, thesensor may be covered with a biocompatible outer membrane that alsoinhibits exposure of the inner membranes to proteins or other largemolecules that may alter the properties of the sensor inner membranes.Such a membrane could be composed of porouspolyhydroxyethyl-methacrylate, polyethylene- or polycarbonate-containingpolymers, fluorinated polymers, or other suitable materials.

Desirable sizes and shapes of hydrophilic regions and associatedmembranes can be calculated by a systematic, computational approach. Ina preferred embodiment, the device contains at least one hydrophilicregion over a single electrode (FIGS. 1-4). The sensor is a discplatinum oxygen electrode closely apposed to a hydrophilic region andthe hydrophilic region is surrounded by a material that is essentiallyimpermeable to glucose. The hydrophilic region contains immobilizedglucose oxidase and optionally, an excess of catalase. For a givenglucose concentration in the external medium the sensor response isdetermined by the permeability of the hydrophilic region and membranebody, the enzyme activity, and the aspect ratio, or ratio of the averageequivalent radius of the critical zone within the hydrophilic region tothe height of the critical zone. In order to obtain a useful range ofresponse in biological operating conditions, it is preferred that thisaspect ratio be less than one.

EXAMPLE 1

Sensor membranes were produced by filling the cavities in perforatedsilicone rubber sheets with a glucose oxidase/albumin mixture andcrosslinking the mixture with glutaraldehyde using the method describedin Armour et al. 1990, incorporated herein by reference.

The membranes were mounted over a membrane-covered electrochemicaloxygen sensor, with a circular platinum working electrode of diameter0.005″, formed on an alumina ceramic substrate using conventionalthick-film methods. The required counter electrode was platinum and therequired reference electrode was silver-plated platinum.

The devices were connected to a potentiostat circuit, and the workingelectrode was polarized at −500 mV with respect to the referenceelectrode. (see for example: Bard and Faulkner, 2000).

Tests were conducted in a simulated biological environment:phosphate-buffered saline, at 37° C., equilibrated with known oxygenconcentrations. Known quantities of glucose were added to the solutionand the electrode current measured. Two different membrane geometries,schematically represented in FIG. 5, with the specifications shownbelow, were tested. As is well-known (see e.g. Gough et al, 1985), thedevice's response is suitably analyzed by examination of the normalizedelectrode current as a function of the glucose-to-oxygen ratio in theenvironment. Both raw i(nanoampere) electrode currents and normalizedcurrents (expressed as a percentage of the value without glucose) arereported below.

Specifications

membrane thickness: 0.010″ hydrophilic region shape: funnel hydrophilicregion radius at base (closest to electrode): 0.014″ hydrophilic regionraduis at top, communicating with fluid: 0.003″

Results

[glucose]/[oxygen] electrode current electrode current (mg/dl/mmHg)(nanoamperes) (% of initial) 0 12.8 100 0.98 10.6 83 2.7 9.3 73 5.9 7.559 10.8 5.7 45 22.1 1.0 8

Specifications

membrane thickness: 0.010″ hydrophilic region shape: funnel hydrophilicregion radius at base (closest to electrode): 0.014″ hydrophilic regionradius at top, communicating with fluid: 0.002″

Results

[glucose]/[oxygen] electrode current electrode current (mg/dl/mmHg)(nanoamperes) (% of initial) 0 9.9 100 1.2 8.9 90 2.8 8.3 84 5.7 7.4 7511.1 6.3 64 22.7 3.5 35 42.8 1.2 12

EXAMPLE 2

Sensor membranes were produced by filling the cavities in perforatedsilicone rubber sheets with a glucose oxidase/albumin mixture andcrosslinking the mixture with glutaraldehyde using the method describedin Armour et al. 1990, incorporated herein by reference. The membraneswere mounted over a membrane-covered electrochemical oxygen sensor, witha rectangular platinum working electrode of dimensions 0.025″(inches)×0.2″, formed on an alumina ceramic substrate using conventionalthick-film methods. The required counter electrode was platinum and therequired reference electrode was silver-plated platinum.

The devices were connected to a potentiostat circuit, and the workingelectrode was polarized at −500 mV with respect to the referenceelectrode, following well-known methods (see for example: Bard andFaulkner, 2000).

Tests were conducted in a simulated biological environment:phosphate-buffered saline, at 37° C., equilibrated with known oxygenconcentrations. Known quantities of glucose were added to the solutionand the electrode current measured. Two different membrane geometries,schematically represented in FIG. 6, with the specifications shownbelow, were tested. As is well-known (see e.g. Gough et al., 1985), thedevice's response is suitably analyzed by examination of the normalizedelectrode current as a function of the glucose-to-oxygen ratio in theenvironment. Both raw (nanoampere) electrode currents and normalizedcurrents (expressed as a percentage of the value without glucose) arereported below.

Specifications

membrane thickness: 0.010″ hydrophilic region shape: cylindricalhydrophilic region radius: 0.005″ hydrophilic region spacing: 0.020″center-to-center, offset grid pattern

Results

[glucose]/[oxygen] electrode current electrode current (mg/dl/mmHg)(nanoamperes) (% of initial) 0 74 100 0.6 54 73 1.1 45 61 2.2 36 49 2.831 42 4.2 26 35 5.6 22 30 11.2 14 19 22.4 8 11 44.9 5 7

Specifications

membrane thickness: 0.010″ hydrophilic region shape: cylindricalhydrophilic region radius: 0.005″ hydrophilic region spacing: 0.010″center-to-center, offset grid pattern

Results

[glucose]/[oxygen] electrode current electrode current (mg/dl/mmHg)(nanoamperes) (% of initial) 0 192 100 0.6 123 64 1.1 89 46 2.2 53 282.8 45 23 4.2 31 16 5.6 35 13 11.2 6 3 22.4 2 1

EXAMPLE 3

Optimization of hydrophilic region shape and size was carried out usingcomputer modeling methods. The analysis is based on the modeling ofdiffusion and reaction of glucose and oxygen in the presence of glucoseoxidase and catalase within the hydrophilic region. The chemicalreaction can be summarized as follows:

glucose+½O₂→gluconic acid

Computer models of operating devices were constructed using conventionalmethods (see for example: Jablecki and Gough, 2000, incorporated hereinby reference) to calculate the response of an oxygen sensor, incommunication with one or more hydrophilic regions, to environmentalglucose and oxygen concentrations for various membrane constructions. Inthese analyses, the electrode current is calculated and shown asi_(g)/I_(o), which is the ratio of the glucose-modulated oxygen currentto the current in the absence of glucose (see e.g. Armour, et al 1990).This normalized current equals zero in the absence of glucose and risesto a maximum value of unity as glucose concentration increases.

In all cases, useful sensitivities for monitoring glucose in biologicalmedia are obtainable only if the average equivalent radius of thehydrophilic region's critical zone is less than the length of thecritical zone. If the average equivalent radius is greater than thelength, then the critical zone is not adequately supplied withcoreactant and the device's dynamic response range is too limited forpractical use in biological samples.

The response range and sensitivities were modeled for three differentshapes of hydrophilic regions analogous to those shown in FIGS. 4, 1 and3, respectively. The data demonstrate that parameters may be readilymodified by altering the shape of the hydrophilic region depending onother device considerations well known to those skilled in the art.

FIG. 6A shows the calculated response of an oxygen sensor (radius 62.5microns) in communication with a membrane containing a cylindricalhydrophilic region, of length 350 microns, for various cylinder radii R.In all cases, the cylinder radius is less than the length, and themodeled devices demonstrate acceptable response to glucose.

FIG. 6B shows the calculated response of an oxygen sensor (radius 62.5microns) in communication with a membrane containing a conicalhydrophilic region, with a base radius R2 equal to 250 microns, andvarious values of top radii R1. The cone base is oriented toward theoxygen sensor and the length is 350 microns. In all cases, the averageequivalent radius of the hydrophilic region is less than the length, andthe modeled devices demonstrate acceptable response to glucose.

FIG. 6C shows the calculated response of an oxygen sensor (radius 62.5microns) in communication with a membrane containing the inverted“T”-shaped cross-section hydrophilic region that is depictedschematically in FIG. 3, with a hydrophilic region base radius equal to250 microns, and a top radius R equal to 62.5 microns. The total lengthof the “T” is 250 microns and the responses of the sensor for variouslengths L1 of the small radius section are shown for L1=0 to L1=250microns. Note that for critical zone aspect ratios of radius-to-lengthgreater than 1, the dynamic range of the device is too limited for usein many biological or physiological media. In all cases when the averageequivalent radius of the hydrophilic region critical zone is less thanthe length, the modeled devices demonstrate an acceptable range ofresponse to glucose.

In the optimization calculation, circular cross-sections are used todetermine the preferred size of the hydrophilic regions. However, thisdoes not limit the instant invention to the use of round hydrophilicregions. The optimization calculation provides ideal internal andexternal surface areas and spacing for the hydrophilic regions that maybe any shape. The selection of shape is a matter of choice to be madebased on any of a number of factors including the shape of theelectrodes, the overall shape of the sensor and the ease of manufacture.

Although an exemplary embodiment of the invention has been describedabove by way of example only, it will be understood by those skilled inthe field that modifications may be made to the disclosed embodimentwithout departing from the scope of the invention, which is defined bythe appended claims.

REFERENCES

Armour, J. C., J. Y. Lucisano, B. D. McKean and D. D. Gough “Applicationof a Chronic Intravascular Blood Glucose Sensor in Dogs,” Diabetes39:1519-26 (1990).

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I claim:
 1. A measuring device for determining the concentration of afirst molecule in an environment, which first molecule reacts inpresence of a catalyst within the device, with a second molecule to forma third molecule; the measuring device comprising: (a) a membranecomprising: (1) a body permeable to said second molecule ad essentiallyimpermeable to said first molecule; and (2) at least one discretehydrophilic region in communication with the body, which hydrophilicregion contains the catalyst and is permeable to both the first andsecond molecules; (i) wherein the direction of diffusion of the secondmolecule into the body is substantially parallel to the direction ofdiffusion of both the first and second molecules into the hydrophilicregion; (b) at least one critical zone coincident with the hydrophilicregion, or a portion thereof, which critical zone has an averageequivalent radius and a length, (1) wherein the average equivalentradius of the critical zone is less than its length, and (2) whereinfurther an average vector direction of diffusion of the first moleculewithin the critical zone is substantially parallel to an average vectordirection of diffusion of the first molecule within the hydrophilicregion; (c) at least one sensor, having a surface communicating with atleast one hydrophilic region and sensitive to either of the second orthird molecules, wherein the sensor is adapted to produce a signalindicative of the concentration of the second or third molecules in thehydrophilic region; and, (d) a control, responsive to the signalproduced by the sensor, for comparing the signal to a reference todetermine the concentration of said first molecule in the environment.2. The measuring device of claim 1, wherein the device is adapted forimplantation of the membrane and sensor into an individual.
 3. Themeasuring device of claim 2, wherein further the device is responsive tomolecules diffusing from a biological fluid or mammalian tissue.
 4. Themeasuring device of claim 1, wherein the hydrophilic region and criticalzone of the device permit diffusion of glucose as the first molecule. 5.The measuring device of claim 4, wherein the catalyst is glucoseoxidase.
 6. The measuring device of claim 1, wherein the hydrophilicregion and critical zone of the device permit diffusion of lactate asthe first molecule.
 7. The measuring device of claim 6, wherein thecatalyst is lactate oxidase.
 8. The measuring device of claim 1, whereinthe hydrophilic region and critical zone of the device permit diffusionof cholesterol as the first molecule.
 9. The measuring device of claim8, wherein the catalyst is cholesterol oxidase.
 10. The measuring deviceof claim 1, wherein the membrane body, hydrophilic region and criticalzone of the device permit diffusion of oxygen as the second molecule.11. The measuring device of claim 1, wherein the third molecule formedis hydrogen peroxide.
 12. The measuring device of claim 1, wherein themembrane body is selected from the group of molecules consisting ofsilicone-containing, ethylene-containing and propylene-containingpolymers with and without fluorine, silicone rubbers, polyethylene,polypropylene, teflons and polyfluorinated hydrocarbons,poly-methylmethacrylates, poly-carbonates,poly-hydroxyethylmethacrylate, and co-polymers and combinations thereof.13. The measuring device of claim 1, wherein the hydrophilic region isselected from the group of molecules consisting of polyacrylamide gels,gluteraldehyde cross-linked proteins, vinyl pyrollidone, alginates,ethylene oxide, acrylamide, methylacrylic acids,polyhydroxyethyl-methacr-ylate and its derivatives, and co-polymers andcombinations thereof.
 14. The measuring device of claim 1, wherein thehydrophilic region has essentially an identical surface area on theinner and outer faces of the membrane.
 15. The measuring device of claim1, wherein the hydrophilic region has a larger surface area on the innerface of the membrane as compared to the outer face of the membrane. 16.The measuring device of claim 1, wherein the membrane contains aplurality of hydrophilic regions.
 17. The measuring device of claim 16,wherein the plurality of hydrophilic regions vary in size one another.18. The measuring device of claim 1, wherein said average equivalentradius of the critical zone is obtained by dividing the cross-sectionalarea of the critical zone by pi, and then taking a square root of theresulting quantity.
 19. The measuring device of claim 1, wherein asingle hydrophilic region communicates with more than one sensor. 20.The measuring device of claim 1, wherein the base of the hydrophilicregion is nearly identical in area to the area of a sensor communicatingwith the hydrophilic region.
 21. The measuring device of claim 1,wherein the base of the hydrophilic region is different in area than asensor communicating with the hydrophilic region.
 22. The measuringdevice of claim 1, wherein a plurality of hydrophilic regions arepresent in the membrane, and a single sensor communicates with more thanone hydrophilic region.